medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Multiplexed Detection of Sepsis Markers in Whole Blood using
Nanocomposite Coated Electrochemical Sensors
Uroš Zupančič1,2, Pawan Jolly1, Pedro Estrela2, Despina Moschou2, and Donald E. Ingber1,3,4,*
1Wyss Institute for Biologically Inspired Engineering, Harvard University, USA,
2Centre for Biosensors, Bioelectronics and Biodevices (C3Bio) and Department of Electronic &
Electrical Engineering, University of Bath, UK,
3Vascular Biology Program and Department of Surgery, Boston Children's Hospital and Harvard
Medical School, USA
4Harvard John A. Paulson School of Engineering and Applied Sciences, Harvard University,
USA
*Address all correspondence to: Donald E. Ingber, MD, PhD, Wyss Institute at Harvard
University, CLSB5, 3 Blackfan Circle, Boston MA 02115 (ph: 617-432-7044, fax: 617-432-
7828; email: [email protected]
NOTE: This preprint reports new research that has not been certified by peer review and should not be used to guide clinical practice. 1 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
ABSTRACT
Sepsis is a leading cause of mortality worldwide that is difficult to diagnose and manage
because this requires simultaneous analysis of multiple biomarkers. Electrochemical detection
methods could potentially provide a way to accurately quantify multiple sepsis biomarkers in a
multiplexed manner as they have very low limits of detection and require minimal sensor
instrumentation; however, affinity-based electrochemical sensors are usually hampered by
biological fouling. Here we describe development of an electrochemical detection platform that
enables detection of multiple sepsis biomarkers simultaneously by incorporating a recently
developed nanocomposite coating composed of crosslinked bovine serum albumin containing a
network of reduced graphene oxide nanoparticles that prevents biofouling. Using nanocomposite
coated planar gold electrodes, we constructed a procalcitonin sensor and demonstrated sensitive
PCT detection in undiluted serum and clinical samples, as well as excellent correlation with a
conventional ELISA (adjusted r2 = 0.95). Sensors for two additional sepsis biomarkers — C-
reactive protein and pathogen-associated molecular patterns — were developed on the same
multiplexed platform and tested in whole blood. Due to the excellent antifouling properties of the
nanocomposite coating, all three sensors exhibited specific responses within the clinically
significant range without any cross-reactivity in the same channel with low sample volume. This
platform enables sensitive simultaneous electrochemical detection of multiple analytes in human
whole blood, which can be expanded further to any target analyte with an appropriate antibody
pair or capturing probe, and thus, may offer a potentially valuable tool for development of
clinical point-of-care diagnostics.
2 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
GRAPHICAL ABSTRACT
KEYWORDS: Anti-fouling, electrochemical biosensor, multiplexing, point-of-care diagnostics,
sepsis
Specific and sensitive quantification of multiple biomarkers for point-of-care (POC)
diagnostics applications using whole blood samples has been a holy grail in clinical diagnostics
for decades. The greatest need for such devices is the diagnosis of life-threatening medical
conditions, such as sepsis, where every hour of delay in administration of the correct antibiotic
increases the mortality rate by over 7%.1 Unfortunately, no single sepsis biomarker exhibits
appropriate specificity and sensitivity for disease diagnosis, and thus multiplexed detection of
biomarker panels has emerged as the preferred POC diagnostic option.2 Among the variety of
biosensing platforms, electrochemical (EC) biosensors would be preferred to use for this
application due to their potential for optimal detection limits, cost-effective integration with
diagnostic readers, and low power instrumentation requirements.3, 4 However, while individual
biomarker detection has been demonstrated in whole blood using electrochemical systems,5, 6 it
only has been possible to carry out multiplexed detection by prefiltering the blood, capturing and
then releasing the biomarker molecules,7 which adds greatly to the complexity of a potential
3 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
integrated ‘sample in-answer out’ device. Commercial EC POC platforms have been developed
that utilize separate cartridges for different biomarkers (e.g., Abbott’s i-STAT system), but they
do not carry out multiplexed protein detection and quantification cartridge.8 Therefore, there is a
need for a low-cost and straightforward multiplexed EC platform that can rapidly quantify
multiple biomarkers in a complex sample, such as serum, plasma, or whole blood.
One of the reasons behind the limited commercial success of affinity-based EC sensors in
complex biological samples is the high susceptibility of the EC sensing elements to biological
fouling.9 Nonspecific binding can drastically impact resistive and capacitive properties of the
sensing element, hampering electrochemical signal transduction.10 We recently demonstrated
that this issue can be addressed by coating electrodes with a three-dimensional (3D)
nanocomposite containing crosslinked bovine serum albumin (BSA) doped with conducting
nanomaterials, such as gold nanowires (AuNWs) or carbon nanotubes (CNTs), and using
3,3′,5,5′-tetramethylbenzidine (TMB) that precipitates locally on the surface of the electrode to
quantify binding.11, 12 This matrix allowed excellent electrochemical signal transduction between
the gold electrode and the nanocomposite surface, which could be functionalized with antibodies
while maintaining conductive and antifouling surface properties.11, 12 Although the
nanocomposite demonstrated excellent antifouling characteristics by maintaining its properties
after one-month exposure to human plasma, the addition of expensive AuNWs to the matrix can
pose a significant challenge towards commercialization of this technology. Thus, in the present
study, we replaced AuNWs with reduced graphene oxide nanoflakes (rGOx) with known
antifouling and electrochemical attributes, which are also much less expensive.13, 14 By doing so,
the cost of the nanocomposite was reduced by ~99%, which should enable this technology to
have an immediate commercial impact. The EC sensor platform we present here, which is
4 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
enabled by the graphene oxide nanoflakes-based antifouling nanocomposite, can quantify
multiple biomarkers and can be easily incorporated with microfluidics. An antibody-based sensor
for one of the most important sepsis biomarkers, procalcitonin (PCT), is characterized in serum
and whole blood, and the sensor is validated using clinical samples. The broad applicability of
the nanocomposite is further demonstrated by the incorporation of two additional sensors for two
other infection biomarkers that are not based on antibodies as capture probes: C-reactive protein
(CRP),15 which is based on use of a phosphorylcholine (PC) capturing probe, and pathogen-
associated molecular patterns (PAMPs) that are bound by immobilized Fc-Mannose Binding
Lectin (Fc-MBL).16 This is the first time a multiplexed EC sensor for these three different sepsis
biomarkers has been demonstrated in whole blood.
RESULTS AND DISCUSSION
Nanocomposite characterization
Nanocomposite coatings composed of BSA and AuNWs cross-linked with
glutaraldehyde (GA) have been shown to demonstrate excellent antifouling properties,
maintaining high conductivity even after one-month exposure to human plasma and allowing
sensitive detection of interleukin-6 (IL-6).17 To lower the cost of the nanocomposite, AuNW
were replaced with conducting rGOx nanoflakes that were terminated with either a carboxyl or
amine group. We first constructed an EC sensor using this modified nanocomposite coating and
analyzed its ability to detect IL-6 in undiluted serum to compare its sensitivity relative to the
AuNW sensor studied in the past and confirm its functionality. The EC sensor employing the
nanocomposite with carboxyl terminated rGOx and use of precipitating TMB resulted in a limit
of detection (LOD) of 63.1 pgmL-1 for IL-6 in undiluted human serum, while the amine-
5 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
terminated rGOx was even more sensitive (LOD = 13.6 pgmL-1) (Figure S1). Importantly, the
latter result represents a significant improvement relative to our previous described IL-6 sensor
using the AuNW containing composite (LOD = 23 pgmL-1).11 A high concentration of labeling
molecules contributes to high sensitivity, and we found that concentrations of streptavidin-
polymerized HRP conjugate (Strep-PolyHRP) as high as 1 µgmL-1 could be used successfully for
target labeling with minimal non-specific binding (Figure S2).
Covalent interaction between the conducting nanomaterial and the BSA matrix improves
the efficiency of nanomaterial incorporation, allows for high dispensability, and allows the
composite to maintain its structure over time,18 maintaining its antifouling properties. Altering
the terminal groups of rGOx affects its solubility and agglomeration in solution, and hence the
distribution of the nanoflakes within the nanocomposite and its final stability.19 Amine
terminated rGOx nanoflakes allow covalent linking of the nanomaterial within the
nanocomposite through glutaraldehyde pyridine polymers that are also responsible for
crosslinking of the BSA,11 which may contribute to the increased sensitivity we observed with
these coatings.
Field emission electron scanning microscope (FE-SEM) imaging of the GA crosslinked
BSA/rGOx coating deposited on the gold surface revealed that it is a densely packed composite
(Figure 1A). At the same time, use of atomic force microscopy (AFM) to analyze the film’s
porosity revealed higher surface roughness (RMS = 1.290 nm) when compared to gold with no
nanocomposite that underwent the same treatment (RMS = 731.1 pm) (Figure 1B and Figure
S3) confirming the formation of the cross-linked protein-based nanocomposite on the electrode
surface. AFM topography also revealed profiles similar to previously published gold nanowire
BSA composites (Figure 1C and Figure S4).
6 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
To characterize the electrical properties of the nanocomposite, cyclic voltammetry (CV)
in ferri-/ferrocyanide solution was performed (Figure 1D). As demonstrated previously,11 GA
cross-linked BSA results in a porous coating allowing for migration of electroactive species to
the proximity of the electrode. When comparing the obtained current density of cross-linked
BSA composite to a bare gold electrode, we observed 57.3% of the oxidation and 61.6% of the
reduction currents, which increased when we incorporated amine-terminated reduced graphene
oxide (rGOx-NH2) in the nanocomposite (respectively resulting in 74.0% and 73.7% of oxidation
and reduction currents compared to bare gold). This result indicated that simply imbedding
conductive rGOx flakes increase the conductivity of the surface coating; however, when BSA-
rGOx-NH2 composite was covalently cross-linked with GA, the coating formed displayed 90.5%
and 94.7% of oxidation and reduction currents compared to bare gold, respectively. This
BSA/rGOx/GA nanocomposite maintained the same peak-to-peak separation potential (∆Ep) of
0.151 V, which was significantly higher for the BSA/rGOx composite (0.252 V) and for
BSA/GA alone without embedded conducting materials (0.524 V)(Figure 1D).
7 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Figure 1. Characterization of the nanocomposite coating. (A) FE-SEM image of bare gold (Au)
and nanocomposite coated gold surface (BSA-rGOx-GA). (B) AFM topography representation
of bare gold (Au) and nanocomposite coated gold surface (BSA-rGOx-GA). (C) The extracted
profile from the white line numbered 1 in B. (D) Cyclic voltammograms representing oxidation
and reduction of 5mM ferri-/ferrocyanide solution using gold electrodes with various coatings;
BSA crosslinked with glutaraldehyde (BSA-GA), BSA incorporated with rGOx nanoflakes
(BSA-rGOx) and glutaraldehyde crosslinked BSA incorporated with rGOx nanoflakes (BSA-
rGOx-GA). (E) Oxidation currents of ferri-/ferrocyanide couple obtained after exposure of
nanocomposite coated electrodes to buffer (phosphate buffered saline, PBS), human serum,
plasma, and whole blood. Bars represent the average, error bars are standard deviation of the
mean, n = 4. (F) Oxidation and reduction current density of the ferri-/ferrocyanide couple against
the square root of the scan rate, obtained at different stages in sensor preparation. Bare gold,
nanocomposite covered gold (BSA/rGOx/GA) and Fc-Mannose Binding Lectin functionalized
nanocomposite surface (BSA/rGOx/GA – Fc-MBL/BSA).
The antifouling properties of the BSA/rGOx/GA nanocomposite were confirmed when we
exposed the coated electrodes to serum, plasma, and whole blood, which are the most frequently
used biological fluids used in POC testing. No significant changes in current density was
observed after 60 min incubation with these complex fluids indicating minimal fouling occurred
on the electrodes (Figure 1E). An additional benefit of the BSA/rGOx/GA nanocomposite is its
versatile functionality due to the plethora of different amino acids naturally present in the BSA
protein that can be used for chemical conjugation. The abundance of aspartic and glutamic acid
can be explored by activation of carboxyl acid groups through carbodiimide chemistry for
covalent linkage of target capture probes to the surface.
8 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Large probes can affect surface chemistry and decrease the sensitivity of an
electrochemical sensor due to their size and charge screening effect.20, 21 To gain an insight into
the surface chemistry of the BSA/rGOx/GA nanocomposite, the genetically engineered human
opsonin protein, Fc-MBL, was conjugated to the nanocomposite. The remaining activated
carboxyl groups were quenched in tris buffered saline (TBS) and blocked with 0.1% BSA in
TBS (Figure 2A). Increasing the CV scan rate in ferri-/ferrocyanide solution while analyzing the
Fc-MBL-modified nanocomposite coated electrode revealed a linear relationship between the
current density and the square root of the scan rate, indicating diffusion limiting oxidation and
reduction of electroactive species (Figure 1F and Figure S5). Furthermore, the functionalized
nanocomposite showed distinct, but small deviation from the behavior of bare gold or non-
functionalized nanocomposite, indicating surface characteristics can be obtained even after the
conjugation of large proteins, such as Fc-MBL (MW = 90 kD) to the nanocomposite. Previous
studies using alkanethiol22 and polyethylene‐glycol23 based self-assembly monolayers report
considerably higher loss of electroactivity after surface functionalization with capturing probes.
The multiplexed electrochemical sensor platform
We then leveraged the excellent electrochemical properties of the nanocomposite to
construct a sensitive electrochemical ELISA using the coated electrode (Figure 2B). Due to the
ultra-low fouling, non-specific interactions are minimized when samples come in contact with
the surface. The biotinylated detection mAb and a conjugate of Strep-PolyHRP were introduced
in separate steps to allow for greater versatility during assay development. In the final step
precipitating TMB was used, which precipitates locally to the electrode where the enzyme is
present. The sensor output was read by cyclic voltammetry at 1 V/s to ensure complete oxidation
9 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
of the precipitated TMB on the surface. Local precipitation of the TMB enables the construction
of a multiplexed sensing array where multiple electrodes with various capturing probes can be
placed in close proximity. To demonstrate proof-of-principle, we constructed a sensor with 4
gold working electrodes that shared a pseudo reference gold electrode and counter electrode, all
compacted in an area less than 7 x 2.5 mm (Figure 2C).
Figure 2. Electrochemical ELISA. (A) Preparation of the electrochemical sensor. (B) A stepwise
process involving the application of the sample, biotinylated detection antibody, polymerized
conjugate of streptavidin and horseradish peroxidase, and precipitating TMB. Between each step,
an additional washing process is introduced (see Methods). (C) Planar evaporated gold electrode
chip comprising of four individual working electrodes (WE 1 - 4), shared counter electrode (CE)
and shared quasi reference electrode (RE) allowing for four separate sensors to be constructed in
an area smaller than 7 x 2.5 mm.
Procalcitonin detection in complex biological fluids
10 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Procalcitonin, an FDA approved biomarker used for assessing risk of developing sepsis
in critically ill patients24 rises from 50 pg mL-1 to over 10 ng mL-1 in cases of fulminant sepsis.25,
26 An electrochemical sensor for PCT was constructed on top of the nanocomposite using anti-
PCT monoclonal antibody (mAb), as described above. Step by step validation of the assay
construction was confirmed by surface plasmon resonance (SPR) analysis, which confirmed
specific binding of PCT to the capturing mAb and its binding to biotinylated secondary mAb, as
well as Strep-PolyHRP binding to biotin moieties (Figure S6). When the sensor was tested in
buffer (Figure S7), undiluted serum, and 50% diluted whole blood (Figure 3A) using various
PCT concentrations spanning four orders of magnitude, we observed similar behavior in all three
fluids. In undiluted serum, the PCT sensor revealed a large dynamic range (0.09 – 10.24 ng mL-
1) and an LOD of 64.5 pg mL-1. When tested in 50% diluted blood, the sensitivity of the sensor
surprisingly increased (LOD of 24.7 pgmL-1), but this was accompanied by a decrease in a
-1 -1 dynamic range (0.07 – 2.49 ngmL ). The dissociation constant (Kd = 0.45 ± 0.07 ngmL )
-1 measured in diluted blood also remained comparable to the sensor in buffer (Kd = 0.42 ngmL ±
0.07 ngmL-1). With increasing PCT concentration, the area of TMB oxidation peaks increased
(Figure 3B), and two TMB oxidation peaks could be distinguished due to the excess TMB
precipitated on the surface when abundant amounts of enzyme are present. A thick layer of TMB
can have an insulating effect,10 and higher potentials are needed to complete the oxidation of
TMB that is present further away from the surface.
11 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Figure 3. (A) PCT sensor response in undiluted serum (black) and 50% diluted whole blood
(red). Squares represent the average (n = 3) and error bars the standard deviation. Hill fit was
performed to fit the data. Empty circles represent the response obtained using the control
electrode with no capturing mAb. (B) TMB oxidation peaks obtained with a CV from PCT
sensor in whole blood with a blank and increasing PCT concentrations.
In every experiment, one electrode was not modified with PCT mAb, and this was used
as an on-chip negative control. Analysis of the neighboring electrodes revealed that no current
12 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
peaks were observed in the electrode without capturing mAb even at highest PCT concentration
(Figure S8), again demonstrating the potent antifouling properties of the nanocomposite surface
coating and the highly localized precipitation of catalyzed TMB. The main advantage of POC
devices in the field of infectious diseases is the time it takes from the start of the analysis until
the result.27 Importantly, the turnaround time of this prototype EC sensor was 51 min, which
meets the clinical need of diagnosis within the first hour.1
Microfluidic integration of the PCT sensor
To further reduce the analysis time, the EC sensor chip was coupled with a poly-
dimethylsiloxane (PDMS) block containing a microfluidic channel using double sided adhesive
tape, and detection of PCT in serum was performed under flow at 10 µL/min with a sample size
of 50 µL (Figure 4Error! Reference source not found.A). The implementation of this
microfluidic sensor assay reduced the total time of sample and label incubation to 7 min. Again,
the antifouling coating performed well under flow with minimal cross reactivity demonstrated by
the on-chip electrode control (Figure 4Error! Reference source not found.B). These findings
demonstrated the coating can be used with microfluidic systems and that it retains its properties
under conditions involving increased diffusion. The data obtained with this PCT electrochemical
sensor were fitted using a Hill fit equation, which resulted in determination of a Kd of 8.69 ±
3.85 ngmL-1 with a dynamic range of 1.06 – 48.8 ngmL-1.
13 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Figure 4. Microfluidic integration of the PCT sensor. (A) Schematic representation of
microfluidic integration by incorporation of nanocomposite covered planar electrode to a PDMS
with the channel outlined by double sided adhesive, connected to a syringe pump and a
potentiostat through a chip interface. (B) Response curve obtained with 50 µL sample using flow
of 10 µL/min decreasing the analysis time below 10 min. Squares represent the average (n = 3)
and error bars the standard deviation. Empty circles represent the response obtained using the
control electrode with no capturing mAb.
14 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Clinical study
To further validate the PCT sensor we analyzed 21 patient serum samples obtained as part of a
past clinical study on detection of blood stream infections16 using the non-microfluidic integrated
protocol described above (Figure 3). The data obtained were compared to the results acquired
using a conventional plate ELISA (1:10 sample dilution was used due to the smaller linear range
of the plate-based ELISA). As the electrochemical sensor analysis was performed using
undiluted serum where PCT levels can increase up to 100 ng mL-1,28 responses in 9 samples were
out of range of the EC sensor. However, the remaining 12 samples were quantified using both
methods, and concentrations obtained with plate ELISA and electrochemical sensors were
compared. The values obtained with both methods were plotted in a scatter plot (Figure 5A) and
linear regression analysis revealed an adjusted r2 of ~0.95 representing excellent correlation
between the electrochemical PCT sensor and conventional ELISA. There also was excellent
agreement between the two methods in the lower concentration region (below 1 ngmL-1) when a
Bland-Altman plot was used (Figure 5B); this is especially important when the EC sensor is
used for early sepsis diagnosis, as higher PCT levels are corelated to more severe sepsis.29 It is
worth mentioning that higher variability at higher concentrations was observed which is likely
due to saturation of EC sensor.
When grouping the clinical samples into infected and non-infected patient groups, both
the electrochemical sensor and conventional ELISA successfully distinguished between the two
groups with high confidence; however, some positive sepsis samples still exhibited low PCT
levels in both assays (Figure 5C and Figure S9), clearly demonstrating the disadvantage of
using a single biomarker for sepsis diagnosis. This is a common problem with the PCT assay in
clinical settings ,30, 31, including in the past study using the same clinical samples with another
15 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
PCT assay16; the PCT assay is also not highly specific.32 Thus, current recommendations do not
recommend the use of PCT as a single biomarker for prognosis, and clearly state the need to use
multiple clinical biomarkers for sepsis diagnosis.33-35
Figure 5. Evaluation of electrochemical PCT sensor in a clinical study. (A) Electrochemical
(EC) sensor response vs. ELISA response using sepsis patients’ and non-infected patients’
clinical samples. Circles represent the mean and error bars represent the standard deviation, n = 3
(EC sensor) and n = 2 (ELISA). Red line represents the linear fit. (B) Bland-Altman plot
describing the agreement between EC sensor and plate ELISA. (C) Box plot comparing PCT
levels in infected and non-infected patient groups obtained with EC sensor. Full circles represent
the mean of the measured values, lines in a box represent the median, 25th and 75th percentile,
empty square is the mean value, whiskers represent 10th and 90th percentile and *** represents p
< 0.001.
Multiplexed detection of PCT, CRP and PAMPs for sepsis diagnosis
The use of multiple biomarkers for sepsis diagnosis can increase the test accuracy,
however the biomarkers must be chosen carefully.36, 37 The highest value of the biomarker panel
can be achieved by combining assays for indirect infection protein biomarkers, such as PCT and
16 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
CRP with direct infection markers that measure the presence of pathogens in the blood sample.38
This latter goal of direct detection of pathogens in blood is no trivial task, but we have previously
shown that it can be achieved using the Fc-MBL opsonin protein that binds to over 100 different
pathogens of all classes (gram -/+ bacteria, viruses, fungi, parasites) as well as pathogen-
associated molecular patterns39, 40, such as lipopolysaccharide endotoxin and lipoteichoic acid,
which are present on the surfaces of pathogens and released into blood when the pathogens are
killed.39, 40 Importantly, using a Fc-MBL PAMPs capture assay, we have previously shown that
the presence of pathogens can be detected within patient whole blood within 1 hour after
collection with high sensitivity (>81%), specificity (>89%), and diagnostic accuracy (0.87), even
in patients that are blood culture negative.16
We also used human whole blood in the current study because no processing steps are
needed after sample collection, which eliminates the risk of removing biomarkers and potential
pathogens from patient samples. The previously described protocol for PAMPs capture by Fc-
MBL40 was adapted and the sensor was tested in buffer (Figure S10) as well as 50% diluted
whole blood (Figure 6A) using mannan as a standard. The electrochemical sensor in whole
blood exhibited higher currents at low mannan levels than buffer due to the baseline presence of
low levels of PAMPs in blood, although the dissociation constants were similar (9.3 ± 0.9 versus
5.1 ± 3.0 ng mL-1 for buffer and blood, respectively). Similar observations were made by
comparing buffer and blood-based analysis using an Fc-MBL plate-based ELISA (Figure S11),
which again demonstrated successful transfer of the Fc-MBL-based assay to an electrochemical
platform. Standard deviations were higher in blood as expected, which led to lower sensitivity
when compared to sensing in buffer (LOD = 6.0 versus 4.1 ng mL-1); however, the dynamic
range in blood was improved, which is important for biomarker quantification.
17 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Figure 6. PAMPs sensor based on Fc-MBL technology and CRP sensor based on
phosphocholine. (A) PAMPs sensor response in 50% diluted whole blood. Mannan is added to
the matrix to obtain the calibration curve. (B) CRP sensor in 50% diluted whole blood. Squares
represent the average (n = 3) and error bars the standard deviation while empty circles represent
the response obtained using the control electrode with no capturing probe in A and B. Hill fit was
performed to fit the data in both experiments.
To further increase the sensitivity and accuracy of the potential electrochemical sepsis
diagnostic sensor, a CRP assay was constructed and integrated into the same platform with the
PCT and Fc-MBL-based PAMPs sensors. Plasma CRP levels can increase to over 100 µgmL-1 in
septic patients and the recommended cut-off values are highly variable. For example, a cut-off
value for diagnosis of infection adults has been proposed to be 79 µg mL-1,41 while for neonatal
sepsis it is 4.82 µg mL-1.42 Detection of CRP in parallel to PCT is also challenging, as
concentrations of CRP are three to five orders of magnitude higher than PCT. Thus, to perform
measurements of the CRP in whole blood in the physiological range and obtain a calibration
curve, we depleted the CRP content in whole blood from 1.44 µgmL-1 to 0.37 µgmL-1 with the
18 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
use of biotinylated anti-CRP mAbs bound to streptavidin coated magnetic beads, which we used
to capture the CRP protein and thereby reduce its levels in blood samples (Figure S12).
We first tacked this challenge relating to large differences in concentrations of these two
biomarkers in blood by developing a mAb based sensor in which the secondary antibody was
conjugated directly to a single HRP moiety, hence avoiding large signal amplification through
the Strep-polyHRP conjugate that we used for PCT and Fc-MBL (Figure S13). Although smaller
currents were observed, the sensor was already saturated at 0.37 µgmL-1 of CRP, hence an
alternative approach using a low affinity capturing probe was used. Phosphocholine (PC) is
known to bind to CRP 43, 44 in the presence of calcium ions and its dissociation constant is 5 μM
at a physiological Ca2+ concentration (10 mM);45 this is in contrast to most antibodies that have
dissociation constants within the nM range.46 In our sensors, we conjugated a PC derivative, 4-
aminophenyl-phosphorylcholine (APPC) to the nanocomposite through EDC/NHS chemistry.
When this electrochemical sensor was used to detect CRP in buffer (Figure S14), we still
observed a response in the nM range indicating strong signal amplification through the Strep-
-1 -1 PolyHRP part of the assay with LOD at 10 ngmL and a Kd of 55.7 ± 4.4 ngmL . Performing
EC CRP assay in CRP-depleted blood diluted 1:1 with buffer allowed us to obtain a calibration
curve that was significantly shifted when compared to the assay in buffer alone (Figure 6B).
This may be the consequence of CRP binding to phosphocholine groups in blood lipoproteins
and phospholipids in the membrane of blood cells,44 which would compete with the PC probes
on the electrode surface for the CRP binding pocket. Our studies revealed that the LOD of the
-1 -1 final CRP electrochemical assay was 0.492 µgmL , with a Kd = 1.54 ± 0.1 µgmL and a
dynamic range of 0.63 – 3.76 µg mL-1; this shifted the sensor range from ng mL-1 to µg mL-1,
which enabled direct measurement of CRP in blood.
19 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
To demonstrate the three sensors can be multiplexed within a single chip, electrodes were
individually functionalized with each capturing probe. Each analyte was added at the highest
concentration according to its respective calibration curve, and exposed to the sensor in buffer
(Figure 7). Minimal cross-reactivity was observed in three independent experiments for each
condition demonstrating the nanocomposite technology can be used to create multiplexed
electrochemical sensors with multiple antibody pairs, as long as there is no cross-reactivity
between the antibodies.
20 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Figure 7. Cross-reactivity of PAMPs, CRP and PCT sensors on the same chip. Chips comprising
of PCT, PAMPs, CRP and control sensors on WE1, WE2, WE3 and WE4 were exposed to 31.25
ngmL-1of mannan (A), 10 µg/mL-1 of and CRP (B) and 50 ngmL-1 of PCT (C). Bars represent the
mean value and error bars the standard deviation (n = 3). *** represents p<0.001 while **
represents p<0.01 in a two-sample T-test. The schematic below represents the probe on the
sensors surface (PCT mAb, PC, Fc-MBL and BSA for control) with the addition of respected
targets, leading to specific response of the EC sensor. C. mAb refers to capturing mAb, bio-
PCT/CRP D. mAb/MBL refers to biotinylated detection PCT/CRP mAb and biotinylated MBL.
21 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
To confirm parallel detection of the three analytes using these multiplexed
electrochemical sensors, four chips were constructed and tested as follows. The first chip was
exposed to 50% diluted blood (with depleted CRP levels) with no additional analytes; the second
was exposed to mannan; the third to mannan and CRP; and the final chip to mannan, CRP and
PCT, all in whole blood diluted 1:1 with buffer. An increase in current density was observed
where the analytes were added, and the overall signature of each electrochemical sensor changed
based on the addition of its respective analyte (Figure 8). We observed a response in the PAMPs
sensor only with blood, which could be associated with presence of PAMPS in the blood
constituting the baseline PAMPs level (Figure 6a). Addition of mannan, which is the ligand for
Fc-MBL found in PAMPS, only increased the signal of the Fc-MBL based sensor. When CRP
and mannan were added, both the CRP and PAMP sensors responded, and upon addition of all
analytes, all sensors responded accordingly.
Figure 8. Multiplexed detection of PCT, CRP and PAMPs in whole blood. Chips comprised of
three sensors were exposed to blood with lowered CRP levels, blood with spiked mannan (31.25
22 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
ngmL-1), blood sample with mannan and CRP (10 µgmL-1) and a blood sample with mannan,
CRP and PCT (50 ngmL-1).
CONCLUSION
In conclusion, we demonstrated that conductive, amine-terminated rGOx nanoflakes can
be successfully integrated within a previously described BSA-based nanocomposite and cross-
linked with GA to form a porous conductive layer for electrodes exhibiting excellent antifouling
properties. This nanocomposite significantly lowered the cost and improved the sensitivity of IL-
6 sensors compared to a previous sensor coated with the same nanocomposite containing
AuNWs rather than rGOx. The nanocomposite is highly versatile as it can be functionalized with
analyte-capturing probes ranging from antibodies and engineered proteins to small molecules
(e.g., APPC) to construct affinity-based electrochemical sensors for detection of virtually any
molecular analyte. Detection of an important sepsis biomarker (PCT) was demonstrated in
buffer, serum and 50% diluted whole blood, and the sensor was validated in a clinical study
using patient blood specimens. The electrochemical sensor was also integrated into a
microfluidic assay and detection of PCT in undiluted serum was demonstrated in less than 10
minutes indicating the potential value of this technology for POC applications. In addition, we
constructed a PC-based CRP sensor that enabled simultaneous detection of PCT and CRP in the
same sensor chip with no cross-reactivity, even though the first biomarkers is present at ng mL-1
concentrations, while the latter is in the µgmL-1 range. We also constructed an electrochemical
Fc-MBL based PAMPs sensor and demonstrated its utility in whole blood for the first time.
Finally, we multiplexed all three sensors (PCT, CRP, and PAMPs) on the same electrochemical
platform to develop a prototype sepsis diagnostic test which showed no cross-reactivity enabling
23 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
parallel multiplex detection of their respective analytes in whole blood. Due to the use of a
sandwich ELISA and TMB substrate that precipitates locally with high accuracy, sensors for
multiple different proteins can now be constructed in close proximity, which allows for creation
of multiplexed sensor arrays on miniaturized biosensing chips. When coupled with microfluidics,
these novel multiplexed sensors could detect a panel of biomarkers in whole blood in a matter of
minutes, opening the way to low-cost and easy-to-use rapid sepsis diagnostics. As a platform
technology, these nanocomposite-enabled sensors also may revolutionize POC diagnostics for a
broad range of other diseases and conditions.
METHODS AND EXPERIMENTAL SECTION
Chip preparation
Gold chips were prepared using standard photolithography process by depositing 20 nm
of titanium and 150 nm of gold on a glass wafer, as described previously.11 Prior to use, gold
chips were cleaned by 5 min sonication in acetone (Sigma Aldrich, USA, no. 650501) and then
in isopropanol (Sigma Aldrich, USA, no. W292907). To ensure a clean surface, the chips were
then treated with oxygen plasma using a Zepto Diener plasma cleaner (Diener Electronics,
Germany) at 0.5 mbar and 50% power for 2 min.
Nanocomposite preparation
Amine-functional reduced graphene oxide (Sigma Aldrich, USA, no. 805432) was mixed
with 5 mgmL-1 BSA (Sigma Aldrich, USA, no. 05470) in PBS solution (Sigma Aldrich, USA,
no. D8537), ultrasonicated for 1 h using 1 s on/off cycles (Bransonic, CPX 3800), heated at
105oC for 5 min to denature the protein and centrifuged to remove the excess aggregates. The
nanomaterial solution was then mixed with 70% glutaraldehyde (Sigma Aldrich, USA, no.
24 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
G7776) for crosslinking in the ratio of 69:1, deposited on the glass chip with gold electrodes and
incubated in the humidity chamber for 20-24 h to form a conductive nanocomposite.
Nanocomposite imaging
The nanocomposite was prepared on gold coated glass SPR chips (Reichert
Technologies, Germany, no. 13206060). The images of the nanocomposite on gold were taken
using a JEOL JSM-6301F field emission scanning electron microscope at 5 kV acceleration and
50000x magnification. AFM was used to evaluate the surface topography and roughness with a
Digital Instruments Nanoscope IIIA AFM and Gwyddion software was used for image
processing.
Nanocomposite electrochemical characterization
Cyclic voltammograms (CV) characterizing the conductive nanocomposite were obtained
3-/4- in 1xTBST (Tris buffered saline with Tween 20) with 5 mM CaCl2 with 5 mM [Fe(CN)6] at a
100 mVs-1 scan rate using chip-integrated gold counter and gold quasi-reference electrodes.
Nanocomposite antifouling properties were characterized by incubating nanocomposite coated
chips in PBS buffer, human serum (Sigma Aldrich, USA), human plasma (Sigma Aldrich, USA)
and whole blood (Research Blood Components, USA) for 60 min and washed for 5 min with
3-/4- PBS and analyzed by cyclic voltammetry in PBS with 1 M KCl and 5 mM [Fe(CN)6] at 100
mVs-1.To characterize the reversibility of the system, bare gold electrodes, electrodes covered
with nanocomposite and electrodes with functionalized nanocomposite (i.e. with covalent
attachment of PAMPs probe FcMBL and surface blocking with BSA) were analyzed by CV in
3-/4- 1xTBST with 5 mM CaCl2 with 5mM [Fe(CN)6] by increasing the scan rate from 10 to 1000
mVs-1.
25 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
Preparation of PCT, CRP and PAMPs sensors
After nanocomposite preparation, gold chips were washed in PBS by agitation (500 rpm)
for 30 min and dried with pressurized air. 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide
hydrochloride (Thermo Fisher Scientific, USA, no. 22980) and N-hydroxysuccinimide (Sigma
Aldrich, USA, no. 130672) were dissolved in 50 mM MES buffer (pH 6.2) at 400 mM and 200
mM, respectively and deposited on nanocomposite covered gold chips for 40 min. After surface
activation, chips were quickly rinsed with ultra-pure water, dried and the capturing probe was
spotted on top of the working electrode area. For PCT sensor, PCT capture mAb (Abcam, UK,
no. ab222276) was diluted to 0.1 mgmL-1 in MES buffer. One electrode was always spotted with
0.1 mgmL-1 BSA as a negative control and the chips were incubated overnight in a humidity
chamber. After conjugation, chips were washed and quenched in TBS (50 mM Tris) for 1 h and
blocked with 0.1% BSA in TBS containing 10 mM glucose. Chips were left in the blocking
solution prior to use. After this point all washing steps were done with 1 mL of the washing
buffer comprising of TBST with 5 mM CaCl2. Chips were incubated with the target sample for
30 min with agitation to mimic flow. After sample incubation and the wash, 2 µgmL-1 of
biotinylated anti-PCT detection monoclonal antibody diluted in 0.2% bovine gamma globulin in
TBS with 0.05% pluronic acid and 5 mM CaCl2 was added to the chip and incubated for 15 min
with agitation. Biotinylated label was then washed away and Poly-HRP-Streptavidin (Thermo
Fisher Scientific, USA, no. N200) was diluted to 1 µgmL-1 in 0.2% bovine gamma globulin in
TBS-0.05% pluronic acid with 5 mM CaCl2 and incubated on the chip for 5 min. After the wash,
precipitating TMB (Sigma-Aldrich, USA, no. T9455) was incubated on the chip for 1 min and
washed. Final measurement was then performed in PBST using a potentiostat (Autolab
PGSTAT128N, Metrohm; VSP, Bio-Logic) by a CV scan with 1 V/s scan rate between -0.5 and
26 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
0.5 V vs on-chip integrated gold quasi reference electrode. Peak area was calculated using Nova
1.11 software and corrected with geometrical surface area of the electrode.
Antibody-based CRP sensor was constructed by functionalization of the nanocomposite
with 0.1 mgmL-1 of anti-CRP capture mAb (R&D Systems, USA, no. DY1707) MES buffer.
Separately, the same mAb was directly conjugated to HRP using Lightning-Link™ HRP labeling
kit (RnD Systems, USA) and used as a label at 300 ng/mL.
The PC based sensor for CRP was composed by modifying the nanocomposite with 4-
amino-phenyl-phosphorylcholine (APPC) (Toronto Research Chemicals, Canada, no. A626000)
by dilution to 1 mgmL-1 in MES buffer for bioconjugation. Biotinylated CRP mAb (R&D
Systems, USA, no. DY1707) was used as a label at 1 µgmL-1 in both assays. For PAMPs sensor
construction, FcMBL was conjugated to the nanocomposite at concentration of 0.2 mgmL-1.
Human MBL2 (Sino Biological, USA, 10405-HNAS) was biotinylated using EZ-Link™ NHS-
PEG4-Biotin (Thermo Fisher Scientific, USA) and used as a label at 200 ngmL-1. All three
sensors were evaluated in buffer (TBST with 5 mM CaCl2 and 0.1% BSA), undiluted serum and
whole blood where target was added to the whole blood, which was then diluted 1:1 with sample
buffer comprising of 100 mM heparin, 20 mM glucose, and 40 mM calcium chloride in TBST
before being deposited on the sensor chip.
Detection of procalcitonin in a microfluidic set-up
Nanocomposite preparation, antibody conjugation, surface quenching and blocking was
performed as described earlier. After fitting of the PDMS block and tubing (Figure S15) a tube
was sequentially preloaded with Strep-PolyHRP labeling solution and PCT sample in serum
which was mixed with labeling mAb in a 10:1 ratio moments before tube preloading. Preloaded
tube was inserted to the steel microfluidic port and flown through the sensor surface for 5 min.
27 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
This was followed by a preloaded 1 min wash, 1 min incubation with Strep-PolyHRP solution
and another 1min wash. 10 µLmin-1 flow rate was used in all experiments. Preloaded tubing was
removed and TMB was added manually and incubated statically for 1 min before washing and
performing the measurement in PBS.
Clinical samples
Patient blood samples were collected at the Beth Israel Deaconess Medical Center
(BIDMC) Emergency Department as part of a previous clinical study16. The study was approved
by the BIDMC Committee on Clinical Investigations, and patients or legal designees signed a
written informed consent. The study included adult septic and non-infected individuals (≥ 18
years of age) where sepsis patients were defined by meeting two or more criteria for systemic
inflammatory response syndrome (SIRS) and the presence of an infection. The control group was
recruited from the same department where patients did not show signs of infection nor meet the
two or more SIRS criteria. The collection of control samples was approved by the Harvard
University Faculty of Medicine Committee on Human Studies (protocol number M20403-101) in
accordance with all national or local guidelines and regulations. Clinical serum samples were
stored at -20 °C until defrosted on ice and centrifuged to remove larger aggregates. 10 µL of
undiluted clinical samples were deposited on the chip and the assay was performed according to
the above-mentioned protocol. Standard PCT ELISA (Abcam, UK, no. ab222276) was
performed on Nunc™ MaxiSorp™ ELISA plates according to manufacturer using 10 µL of the
sample diluted in 90 µL of 1% BSA in PBS. Welch t-test was used when comparing infected vs.
non-infected sample groups.
28 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
CRP removal from whole blood
100 µL of Pierce™ streptavidin magnetic beads (Thermo Fisher Scientific, USA) was
washed three times using 500 µL 1% BSA in PBS before mixing with 3.6 µg of biotinylated
anti-CRP mAb and incubated for 1h with mixing at RT. After incubation, beads were washed 5
times using 500 µL 1% BSA in PBST and added to 0.5 mL of whole blood. Capturing was
performed at RT for 3 h, before beads were removed using a magnet. CRP removal was
evaluated by centrifugation of blood to obtain plasma which was then analyzed by CRP DuoSet
ELISA (R&D Systems, USA, no. DY1707) according to manufacturer.
Multiplexing detection in whole blood
Cross-reactivity of the three sensors was tested by modifying individual electrodes on a
single chip with different capturing probes (PCT mAb, APPC and FcMBL protein). PCT (50
ng/ml), CRP (10 µg/ml) or mannan (31.25 ng/mL) were prepared in TBST with 5 mM CaCl2 and
0.1% BSA. 9 chips were prepared in total and 3 were incubated with PCT, 3 with CRP and 3
with mannan. Different biotinylated labels (bio-PCT mAb, bio-CRP mAb and bio-MBL) were
mixed and used as a label master mix in all 9 chips to test for cross-reactivity. Multiplexing in
whole blood was performed in the same fashion using a mix of biotinylated labels.
ASSOCIATED CONTENT
Supporting Information.
The following files are available free of charge.
Supplementary material includes comparison of carboxy and amine terminated rGOx on
IL-6 sensor sensitivity, antifouling nanocomposite properties in high concentration of Step-Poly
29 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
HRP, further FE-SEM, AFM and electrochemical characterization of the nanocomposite, sensor
validation using SPR, PCT response curve in buffer, demonstration of cross-talk between
neighboring electrodes, comparison of infected and non-infected clinical samples using ELISA,
PAMPs sensor response in buffer, FcMBL ELISA analysis, mAb based CRP sensor response,
PC based CRP sensor in buffer and microfluidic cell set-up for PCT detection (PDF).
AUTHOR INFORMATION
Corresponding Author
* E-mail: [email protected].
Author Contributions
UZ and PJ conceived the study and performed the experiments under the direction of D.M., P.E.,
and D.E.I. ; UZ analyzed the data, which was reviewed by all the authors. All authors
contributed to the writing of the manuscript, and all approved the final version.
Notes
The authors declare no competing financial interests. PJ and DEI are listed as inventors on
patents describing this technology.
ACKNOWLEDGMENTS
The authors would like to thank Drs. M. Super, M. Cartwright, B. Seiler, S. Rifai and N.
Shapiro for supplying materials and clinical samples. This work was supported by Defense
Advanced Research Projects Agency (DARPA) Contract W911NF-16-C-0050, the Wyss
Institute for Biologically Inspired Engineering at Harvard University and the Rosetrees Trust
30 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
(project M681); the salary of U. Zupančič was provided by University of Bath and Santander.
The project was supported by the U.S. Army Research Office, and the content of the information
does not necessarily reflect the position or the policy of the Government, and no official
endorsement should be inferred.
Abbreviations
APPC, 4-aminophenyl-phosphorylcholine; AFM, atomic force microscopy; AuNW, gold
nanowires; BSA, Bovine serum albumin; CRP, C-reactive protein; CV, cyclic voltammetry; EC,
electrochemical; EDC, 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide; ELISA, enzyme-linked
immunosorbent assay; FDA, Food and Drug Administration; FE-SEM, field-emission electron
scanning microscope; GA, glutaraldehyde; IL-6, interleukine-6; LOD, limit of detection; MBL,
mannose binding lectin; NHS, N-Hydroxysuccinimide; PAMPs, pathogen associated molecular
patterns; PBS, phosphate buffer saline; PBST, phosphate buffer saline with 0.5% tween 20; PC,
phosphocholine; PCT, procalcitonin; PDMS, polydimethylsiloxane; POC, point-of-care; rGOx,
reduced graphene oxide; rGOx-NH2, reduced graphene oxide with amine functionalization;
RMS, room mean square; RT, room temperature; SPR, surface plasmon resonance; TBS, Tris-
buffered saline; TBST, Tris-buffered saline with 0.5% tween 20; TMB, 3,3',5,5'-
Tetramethylbenzidine; WE, working electrode.
31 medRxiv preprint doi: https://doi.org/10.1101/2020.11.03.20224683; this version posted November 4, 2020. The copyright holder for this preprint (which was not certified by peer review) is the author/funder, who has granted medRxiv a license to display the preprint in perpetuity. All rights reserved. No reuse allowed without permission.
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